This invention relates to a magnet assembly for a magnetic resonance imaging system (hereinafter called "MRI"), and more particularly to an improved and simplified arrangement for improving magnetic field homogeneity in a plurality of related magnets, including those with small imaging volumes and stringent homogeneity requirements which can not be accurately mapped because of measurement equipment limitations.
As is well known, a magnet can be made superconducting by placing it in an extremely cold environment, such as by enclosing it in a cryostat or pressure vessel containing liquid helium or other cryogen. The extreme cold ensures that the magnet coils are made superconducting, such that when a power source is initially connected to the coils for a short period to introduce a current flow through the coils, the current continues to flow even after power is removed due to the absence of resistance, thereby maintaining a strong magnetic field. Superconducting magnets find wide application in the field of MRI.
However, MRI requires very strong or large magnetic fields in a small imaging bore or volume with a very high degree of uniformity or homogeneity. Typically this requirement which has been on the order of 10 parts per million (ppm) inhomogeneity on a 45 to 50 cm diameter spherical volume (DSV) cannot be achieved by controlling manufacturing tolerances. After manufacturing the magnet with the best achievable tolerances, the inohomogeneity is typically one or two orders of magnitude above the desired level, and a magnetic field shimming system is used to reduce the inhomogeneity level from several hundred ppm to the desired 10 ppm over the 45 to 50 cm spherical volume.
While the problem of shimming a magnet to 10 ppm on a 45 cm volume has been achieved, more demanding applications require smaller levels of inhomogeneity on smaller volumes. These demanding applications include cardiac imaging and high speed MRI scanning using Diffusion Weighted Echo Planar Imaging (DWEPI). Such scan systems are most sensitive to inhomogeneity. However, present test equipment does not enable practical magnetic field mapping at such low inhomogeneity levels. These applications require reduced inhomogeneity levels such as in the order of 0.5 ppm on 25 cm DSV, and 1 ppm on 35 cm DSV to enable the effective use of MRI.
In practice, shim systems utilize extra coils, typically called correction or shimming coils, small pieces of iron, typically called passive shims, or some combination of the two to correct or improve the magnetic field homogeneity while allowing reasonable manufacturing tolerances. Current flow provided through the shimming coils produce magnetic fields to cancel and/or minimize the magnetic field inhomogeneities in the imaging volume.
Considerable engineering and development effort has been applied for some time to improving and simplifying, to the extent possible, the magnetic field homogeneity systems for MRI devices in order to improve imaging quality with uncomplex means and without additional expense. In addition, it is desirable to utilize existing shimming hardware.
Thus, there is a particular need for an MRI shimming system which provides improved magnetic field imaging homogeneity for small volumes and yet which is uncomplex, minimizes the shims used, and utilizes to the extent possible existing test equipment and shimming hardware.